Biomedical Materials and Diagnostic Devices

PART I

BIOMEDICAL MATERIALS

Chapter 1

Application of the Collagen as Biomaterials

Kwangwoo Nam and Akio Kishida

Institute of Biomaterials and Bioengineering, Tokyo Medical and Dental University, Tokyo, Japan

Abstract

Collagen is the protein of connective tissue in mammals. The content of collagen in the total protein is approximately 30% of the mammalian tissues. Due to its good cytocompatibility, researchers use this material for the biomedical research application. However, the control of its physical and biological properties is difficult. There are two obstacles in collagen application: 1) difficulty in regeneration of the collagen properties, and 2) difficulties in controlling the properties of the collagen products. The collagen is easily denatured and affected by the environment, which leads to unexpected results. On the other hand, the crosslinker to suppress the denaturation may cause the stiffness of the collagen product. So the researchers are investigating new ways to prepare a collagen product which can be used as a biomaterial for biomedical research application. An important component of the research is the structure and the function of extracelluar matrix (ECM). That is, there is biorelevant structure-function-property relationship, which alters its function as an ECM. Recent studies on decellularized tissue is also based on the fact that the native structure of the ECM can be preserved, and therefore may perform the function of the original tissue. So, by replicating its microstrutcure and producing a collagen fiber complex, it is expected that the function of ECM can be replicated. In this chapter, we will be introducing recent studies on the preparation of a collagen matrix based on fibrillogenesis, orientation, complex formation and layered structure, and how these structures alter the physical and biological properties.

Keywords: Collagen, decellularization, extracellular matrix, fibrillogenesis, microenvironment, regenerative medicine

1.1 Introduction

Collagen is an extracellular-matrix (ECM) protein that plays an important role in the formation of tissues and organs and is involved in various functional expressions of cells [1]. A native ECM is a complex fiber-composite material in which collagen fibrils are a major component [2]. The function of an ECM is to provide support, tensile strength, and scaffolding for the tissue and cells. In addition, it should serve as a three-dimensional structure for cell adhesion and movement and as a storage depot for growth factors, chemokines, and cytokines; and it should provide signals for morphogenesis and differentiation [3]. Approximately 30% of all vertebrate body protein is composed of collagen. Among these, the highest collagen composition can be found for the tendon, bone and cornea where 90% of ECM is collagen. Mainly, the collagen can be distinguished into two types; fibrillar and non-fibrillar. There are 28 types of collagen and the collagen types I, II and III are the classical fibril-forming collagens and account for 80–90% of all collagens in the human body. Collagen fibril is very important from the aspect that its properties and the morphology provide the key to the scaffolding structures in the body according to the location.

It has been shown that the collagen possesses non-immunogenicity and good cell compatibility, and can be obtained from various sources. These make collagen popular among biomaterials researchers, and diverse methods have been adopted for its application in the biomedical fields. The collagen is purified after being treated with pH adjustment or pepsin digestion. Either way, the collagen should be water soluble in order to process it for use as a collagen matrix for biomaterial applications. There are several kinds of collagen matrix; gel, film, micropartices, conjugats, minipellets or sponge [1, 4]. However, there are still many problems to overcome. For example, the collagen which is available in the marketplace is hydrophilic, which absorbs water at a high rate. So, the uncross-linked collagen matrix possesses low mechanical strength and fast degradation rate in aqueous solution. The collagen matrix degrades by the collagenase, so this makes the collagen applicable in some biomedical products where the biodegradation in the living body is required. However, control of the biodegradation is not easy. The properties of the collagen matrix can be controlled by cross-linking. The cross-linking is executed chemically or physically. Furthermore, using the same cross-linking process, the collagen matrix can be functionalized by immobilization or, blend of a second component. The collagen is composed of amino acid groups where the chemical reaction can be executed. Mainly, the cross-linking is executed using ε-amino groups of lysine or hydrolysine, and aspartic acid or glutamic acid residues. These residues are highly reactive and can be easily functionalized. The cross-linking can change physical and biological properties of the collagen matrix and can be applied for the loading of the drugs. For the chemical cross-link, glutaraldehyde, formaldehyde, hexamethyelenediisocynate, polyepoxy compounds, carbodiimides, and acyl azides are commonly used [1, 4–14]. These show a good result in vivo, such as suppressing the inflammatory response and promoting the healing response. However, there are still several problems to be overcome. Although the collagen gels, sponges or films that have been cross-linked show an increase in the mechanical strength, the cross-link which consumes the functional groups are consumed for the cross-linking site, which may affect the biological properties. Moreover, the stiffness of the ECM is also a very important parameter, but it is not easy to control the stiffness of the collagen gel by cross-linking or a change in the collagen solution. That is, a stiff collagen gel can be prepared, but a gel with viscoelasticity cannot be prepared. This is important because most of the native tissue possesses visocelasticity which contributes to the toughness [15, 16]. A number of collagen matrices were reported, but the collagen products available in the marketplace is still scarce because of the problems mentioned above, so a new approach was required to move to the next step.

1.2 Structural Aspect of Native Tissue

1.2.1 Microenvironment

Before designing a biomaterial, it is necessary to understand the environment of the living body. When the biomaterial is designed for application in tissue engineering and regenerative medicine, the objective is the repair and remodeling of the damaged ECM and tissue, ultimately regenerating its function. The function of the ECMs is deeply related to the behavior of the cells which is affected by the cell-materials interaction. That is, the control of the cells behavior is very important in the aspect of regenerating the function of the ECMs. The cell is immersed in a dynamic landscape composed of insoluble macromolecules of the ECM, soluble bioactive factors and neighboring cells [17]. The environment which controls the fate of the cell inside the living body is called cellular microenvironment. It is very important in the aspect that the ultimate tissue structure and its function are decided by factors contributing to the cellular microenvironment. For this, there needs to be a fundamental understanding on the cellular microenvironment for the materials design. The cellular microenvironment is the environment in the living body which controls the fate of the cells. The microenvironment is composed of signals from the neighboring cells, physical stimuli, soluble factors such as growth factors, and insoluble factors such as ECM. The ECM has been shown to influence cell mitogenesis and chemotaxis [18, 19], direct cell differentiation [20–23], and to induce constructive host tissue remodeling responses [24–26]. The cells from the ECM sense, integrate and proceed the signals to determine behavior and functions, and the information is passed bidirectionally as the microenvironment is remodeled by the cells.

Development of biomaterials for tissue engineering and regenerative medicine has been approached mainly from the aspect of controlling the soluble and insoluble factors. As for the insoluble factor, diverse materials – natural or synthetic – are being investigated. The main goal of using these materials is to replicate the function of the ECM temporarily or permanently. By loading soluble factors in the materials, researchers tried to control the fate of the cells or stimulate the regeneration of the damaged tissues. On the other hand, manipulation of the morphology, microphase, surface physical properties and chemical properties of the material is a major approach for the control of the insoluble factors. These methods show good results and some of them are actually used for clinical practice.

Since the ECM is mainly composed of collagen, use of collagen to replicate its function is actively executed. It should be noted that the function of the ECM is different according to the type of tissue such as cornea, brain, skin, tendon, or blood vessel, where they need to perform a certain function. So the design of the ECM using collagen should be different according to the targeted tissue. However, although the tissues perform different functions, the common aspect of the tissue is that all are made up of collagen fibrils. That is, in order to design a material which may replicate the function of ECMs, fibrillized structure should be considered. Furthermore, it should be acknowledged that the design should include a nanometer to centimeter scale. The schematic structural images of respective tissue from the nanometer to centimeter scale are shown in Figure 1.1. Yip discussed the importance of careful consideration of biorelevant structure-function-property relationships in the design of biomaterials [27]. That is, the regeneration of the physical properties of native ECM is important for the regeneration of biological properties. The importance of the structure can be seen in research related to the decellularized tissue which is discussed in the next section.

Figure 1.1 The scale from the nano-scale to the macro-scale. The collagen matricesdconsist of nano-scale no larger than the microfibril, but the actual scale of the ECM or the tissue is much larger.

1.2.2 Decellularization

The decellularized tissue is a native tissue in which the cells are eliminated by certain treatment. Decellularization of tissue is based on the fact that preservation of the native ultrastructure and composition of ECM is possible [26]. The methods for the decellularization include use of chemical agents (ionic detergents, non-ionic detergents, acids and bases, hypotonic and hypetonic solution, and solvents), biological agents (enzymes and chelating agents), and physical treatment (temperature, pressure and electroporation). It should be understood that every cell removal agent and method will alter ECM composition and cause some degree of ultrastructure disruption. For example, the use of some chemical agents such as sodium dodecyl sulfate (SDS) may cleave the collagen fibrils, but use of physical treatment such as high pressurization would not affect the main structure [26, 28–30]. Furthermore, incomplete rinsing of chemical agents or the cell debris after decellularization process may cause toxicity. However, the minimization of these undesirable effects, rather than complete avoidance by the living body, is the objective of decellularization. So, the focus is set on complete removal of the cells and preservation of the ultrastructure. The methods for the decellularization should be carefully considered according to the density of the fibers, the thickness and the lipid contents. Moreover, the complete washing of the cell debris or the chemical agents after decellularization should be executed because this could cause toxicity.

The mechanical strength after the elimination of the cells is maintained and the regeneration around the implanted decellualrized tissue occurs without serious inflammatory response. So, the decellularization can be executed for the partial or full organs. It should be noted that the native tissue possesses complex structure and the whole structure – either macro or micro – is maintained after the appropriate decellularization process. Furthermore, the degradation of the decellularized tissue is slow, and the remodeling of the damaged tissue occurs without any problems. The regeneration within the living body occurs on the implanted decellularized tissue and starts to function as a replacement. Furthermore, the high mechanical strength of the decellularized tissue would endure the physical stress inside the living body [29, 31–32]. So many decellularized tissue products such as dermis, heart valve, blood vessel, bone and so on, have been introduced to the markets and are enjoying success.

1.2.3 Strategy for Designing Collagen-based Biomaterials

The key for the success of the biomaterials for regenerative medicine is control of the cells’ fate which depends on the materials characteristics; three-dimensional ultrastructure, surface topology and composition of the ECM [17]. The successful point for decellularized tissue is that the three-dimensional ultrastructure, surface topology and composition of the ECM is maintained after the process. So, in order to reproduce the physical and biological properties of the ECM, we should first mimic its three factors as written above. The key points are the fibril formation, orientation, complex formation with second component such as GAG or elastin, and multiple layers. Since the structure of ECM differs according to the tissue, the mimicking of the structure should also be different according to what kind of tissue the researchers want to make. This is because the key function is different according to the tissue. For example, the tendon should have fibrillar structure with high orientation, the cornea should have fibrillar lattice structure, blood vessels should possess elastin-complex fibril structure with multiple layers and high orientation of collagen fibers, and skin should have elastin-complex fibril structure disregarding the orientation. Such ECM structures allow the various tissues to possess certain physical and biological properties adequate for functional performance. So, the structural consideration for replicating the function of tissue is very important.

Many articles consider this point and try to create an ECM resembling collagen structure. The ECM structure consists of collagen fibrils with a second component such as GAG or elastin forming collagen fiber complex. The collagen fiber complex is usually aligned and multi-layered. Such structure is very important from the aspect not only of the fate of the cells, but also for mechanical endurance performance against the stress given by the living body. This distinguishes the ECM from the other monolithic structure where single performance is expected. In most cases, a gain of physical properties would result in a loss of desired biological properties. The ECM possesses highly complex structure which is not easy to replicate. However, many trials for the creation of complex structure with controlled physical and biological properties that have been reported on during past decades mainly focused on controlling the ultrastructure, surface topology and composition of the collagen matrix. The next section is divided into four parts: preparation of collagen matrix based on fibrillogenesis, orientation, complex formation and layered structure. We will be discussing the most recent methods for the preparation of collagen matrix focused on these four subjects.

1.3 Processing of Collagen Matrix

1.3.1 Fibrillogenesis

The change of the ionic strength, pH, or hdrophobicity by additives in the aqueous solution may drive the alteration of collagen molecule alignment with certain regularity as shown in Figure 1.2. The alignment of the collagen molecules results in the formation of the fibrils which causes the precipitation (Fig. 1.2b). This is called fibrillogenesis or collagen reconstitution. The Fibrillogenesis is an aggregation of the collagen molecules which is an entropy driven process. The loss of solvent molecules from the surface of protein molecules results in assemblies with a circular cross-section, which minimize the surface area against the volume ratio of final assembly [33]. Hydrophobic residues of collagen (Leu, Ile, Val, Phe and Trp) play the main role in lateral aggregation [34, 35]. The fibrillogenesis occurs in an aqueous condition with a certain amount of salt. It is thought that formation of salt bridges by the salt is also a major driving force for the formation of fibril with certain periodicity [36]. However, it is also argued collagen fibrillogenesis is driven primarily by the formation of hydrogen-bonded water clusters bridging recognition sites on opposing helices, and that and hydrophobic interactions between opposing non-polar amino acid side chains is not a major driving force of collagen self-assembly [37]. However, the fact is that the physiological ionic strength and neutral pH and increasing temperature, induces spontaneous assembly of type I collagen into native-like fibers and hydrophobic interactions, salt bridge and hydrogen-bonded water clusters cannot be ignored.

Figure 1.2 The schematic image of fibrillization according to the additives (a) and the photographic images of fibrillogensis in NaCl aqueous solution and in heparin aqueous solution (b).

The preparation of the collagen matrix is based on mixing the collagen aqueous solution with a certain amount of salt to adjust the physiological condition. The mixture depends on how the collagen matrix is going to be used. For example, when the collagen and cell are mixed together, cell culture medium can be directly put into the container. The resulting material is mainly gel or sponge which has random fibrillized structure. The temperature for the fibrillogenesis is generally 37°C, but diverse temperature can be used if kept under the denaturation temperature. The thickness of collagen fibers can be controlled where the thickness increases in lower temperature [33]. The drawback of producing a gel or sponge with fibrils at 37°C is that the collagen may denature [38, 39]. The cross-linking usually executed for the gel or sponge is mechanically too weak to support cell growth, proliferation and migration. The most usual phenomenon is the contraction of the matrix caused by the strong cell and matrix surface interaction. However, the cross-linking makes the matrix too brittle [10, 40]. Furthermore, the cross-linked collagen matrix does not show, or at least shows very slow degradation by collagenase [41, 42], which also implies that the possible capsulation in vivo might occur. However, a matrix which does not show contraction upon cell culture and slow degradation in vivo despite the unuse of cross-linker was successfully created. By trapping the collagen aqueous solution in the dialysis cassette and letting the NaCl diffuse into the collagen dialysis, the collagen aqueous solution turned into fibrillzed gel [43]. This gel can be processed into a thin membrane, where the collagen membrane showed a much tougher mechanical strength and slower biodegradation rate in vivo. What is interesting is that this matrix showed suppressed inflammatory response in vivo. The behavior was almost the same as that of decellularized tissue which makes it possible to claim that the structural aspect is very important.

1.3.2 Orientation

The fibril formation can be obtained by the methods described earlier. But the orientation of the collagen fibrils is another problem. The orientation of the collagen molecules or fibrils is reported to be achieved by applying certain force. That is, if there is a certain driving force which allows the collagen fibrils to align, the collagen matrix with orientation can be obtained. The most commonly used method is flow chamber, which allows the collagen molecules to flow into the chamber and precipitate along the axis of the flow. Lanfer et al., have reported on how shear flow deposition would affect the orientation and density of the fibrils [44]. They have concluded that the degree of collagen fibril orientation increased with increasing flow rates of the solution, while the matrix density increased at higher collagen solution concentrations. The rearrangement of the collagen molecules is also reported to be achieved by controlling the concentration of the collagen solution. This is because the collagen behaves much like a liquid crystal and tends to reorganize in high concentration. The Giraud-Guille research group defined the collagen molecules as spontaneously self-organizing in vitro as cholesteric liquid crystal [45, 46]. The fibrillogenesis makes the liquid crystalline phase stable, inducing sol-gel transition. This alignment is especially advantageous for dense collagen, where the direct application for the dermal substitute is expected [47].

An alternative method for collagen alignment is the application of uniaxial elongational strain. This method is advantageous from the aspect that a high orientation percent is obtainable, and direct application on ligaments or tendons is possible. A good example has been reported by Falini et al. They reported on applying uniaxial elongational strain and then dehydrating it for 24 hrs [48]. They concluded that the strain of the collagen film or gel would cause the rearrangement of the collagen molecules when higher than 12% of elongational strain is applied and dehydrated. The orientation percent was approximately 83%, which is very close to the Achilles tendon. However, this method is only applicable to collagen molecules and not to the collagen fibrils. Ross and his group also used an approach for the mechanical strain, but this time they repeated the strain procedure for a longer period of time (2.5% cyclic strain for 2 h per day for 4 days) after seeding the cells [49]. They showed that the alignment of cells along the collagen matrix reflects a response of the cellular environment to the applied strain, concluding that manipulating signal transduction pathways by engineering implantable anterior cruciate ligament grafts or modifying ACL healing response is possible.

Another promising approach is the use of electromagnetic field [50–52]. The collagen molecules tend to align perpendicular to the electromagnetic field upon gelling at above IT. One great advantage of this method is that a high electromagnetic field can be applied to the collagen aqueous solution containing cells and culture condition. The major target is the regeneration of ECM which requires high tensile strength or the neuron tissue [50, 51]. For these, osteoblast, Schwann cell, glioblastoma cell and erythrocyte were cultured upon the collagen gel in which an electromagnetic field above 8T was applied. Y. Eguchi et al. reported that in the mixture of Schwann cells and collagen, Schwann cells oriented in the direction perpendicular to the magnetic field after 2 h of magnetic field exposure. In this case, Schwann cells aligned along the collagen fiber oriented by magnetic fields [52]. This means that the cells and the collagen orientation can be achieved contemporarily. The only problem that remains is how many collagen fibrils would align, because 100% of collagen can not align according to the applied electromagnetic field. Nonetheless, this method remains very promising for tissue engineering.

1.3.3 Complex Formation and Blending

Complex formation of collagen fibrils with a second component is one of the most difficult parts. Complex is mainly executed for the purpose of functionalization. It is generally known that complex formation can be achieved by mixing collagen solution with the second component and cross-link. However, there are some limitations which involve difficulties in collagen molecular control. This is because of the limitation of collagen complex formation with another component. It is known that collagen molecules require hydrogen bond when forming complex with a second component in aqueous solution [53]. Basically, the fibril formation and the complex formation does not occur contemporarily because the second component added to the collagen aqueous solution would function as a defect and prohibit the molecules to aggregate for fibrillogenesis (Figure 1.3). So the complex formation is commonly executed after fibrillogenesis. The most well-known material for collagen complex formation by far is hydroxyapatite, which is designed for bone regeneration. The bone mainly consists of collagen and hydroxyapatite composite where the hydroxyapatite molecules exist between the collagen fibers, providing stiff mechanical strength [54, 55]. The process of collagen-hydroxyapatite composite also targets the creation of such composite structure by diffusing calcium and phosphate ions into the collagen fibrils. The process, also called mineralization, is advantageous in the aspect that the collagen orientation may be achieved at the same time. It should be noted that not all collagen-hydroxyapatite composite is composed of collagen fibrils, but it is still one of the most advanced field.

Figure 1.3 The schematic images of layered structure by layer-by-layer deposition.

Collagen complex is also executed with synthetic polymers, GAG, proteins or oligo-(poly-) saccharides [9, 14, 56–63]. Fibrillized collagen-hydroxyapatite complex involves the diffusion of calcium and phosphate ions. However, this is not easy when the second component is oligomer or polymer. So, the collagen is not necessarily fibrillized through fibrillogenesis for the complex formation in these cases. Instead, the collagen molecules and the second components are mixed together in certain conditions and chemically cross-linked. This method usually involves a freeze-saw process after mixing two solutions or slurry from a sponge, or absorption of polymer solution into the collagen gel or film. These processes allow the formation of large pores suitable for cell migration and develop into the three-dimensional cell culture. Most often hyaluronic acid is used for the collagen complex, for it is known to enhance the cell migration. Furthermore, the existence of hyaluronic acid may induce the moistening affect, which allows the complex matrix to be applicable for the artificial skin. An alternative method for the collagen complex formation is electrospinning. The electrospinning allows the collagen and second component to form fibril blend [61, 64]. The fibrillized structure provides relatively higher mechanical strength than collagen gel or sponge. Furthermore, various polymers can be applied for the blending and can also be produced in a highly aligned state. However, several researchers argue that although electrospinning is advantageous from the aspect that nanometer scale collagen fibrils can be formed, the collagen denatures, making it electrospun gelatin fibrils instead [65]. The importance of collage fibrils possessing regulated D-periodicity is also pointed out by some researchers. This is because the lack of D-periodicity in the electrospun collagen fibrils may cause diseases such as osteogenesis imperfecta and induce cardiovascular disease [2, 66]. The question of using electrospinning for collagen fibrillization is still debated, but diverse blending ability between collagen and the second components, as well as good results of electrospun collagen products in vivo cannot be denied, and many products can be found in the market.

One of the most active research areas is collagen-biodegradable polymer complex. Designed to functionalize the collagen matrix, the polymer eventually degrades together with collagen. The favorite polymers are PLLA and PCL which are known for their nontoxicity and good biodegradability [65, 67–70]. The biodegradability is not a mandatory requirement. This sounds like an oxymoron, but the nondegradable polymers of the collagen complex are usually destined for the high mechanical strength. Sionkowska group uses hydrogen bond inducing polymers such as PVA, PVP or PEG to form complex with collagen molecules [53, 71–73]. Biodegradable polymers are usually applied in electrospinning with collagen in organic solvent. Good blending and altering the fibril diameter can be obtained via this method. So far, the electrospinning process is the only confirmed method for producing collagen fibril-polymer complex. The most interesting blend is a collagen-elastin blend designed for blood vessels which showed very good viscoelasticity [74, 75]. However, it should be noted that this is a blend, not a complex, which implies that there is no chemical or physical interaction between the collagen and the elastin. The neccessity for the collagen to form complex instead of blend remains a question to be answered in the future.

1.3.4 Layered Structure

One of the most important parts of the native tissue is the layer. The idea of multiple layers was brought up for multi-functionalization. For example, the collagen side possessing different functions on each side can be prepared if such a matrix can be prepared. The best way to prepare a multilayered collagen matrix is to adjoin the collagen matrices using adhesives. However, the adhesive would alter the properties of collagen matrices at the interface of the layers and would consist of pure collagen. So a method for the collagen–collagen integration was investigated. The collagen does not integrate with other collagen once they are in solid form. Furthermore, the collagen matrix with fibrils does not normally absorb the polymer into its matrix. As a result, the collagen–collagen interface with entangled layers with polymer as the intermediate do not form. The immobilization technique is the only method that was actually possible for the collagen layer. The immobilization technique involves cross-linking the collagen.

One simple method is the slow drying process. Nam et al. showed that once the collagen matrix with fibrillized structure was formed, the microlayers formed along an axis perpendicular to the surface by the slow evaporation of water [76]. Unlike the lyophilization or dehydrothermal processes, water is not completely eliminated by dehydration. The fibril rearrangement of collagen fibers by dehydration may induce the stability of the collagen matrix against heat as well as its dimensional stability in water. The dehydration causes the chain rearrangement of the collagen molecules, in which the collagen molecules are brought closer to each other [77, 78]. The stripping of the water bridge, which is connected via hydrogen bonds, occurs contemporarily, but the water bridge itself plays a minor role in the stability of the collagen triple helix, indicating that the air drying process does not cause the denaturation of the collagen matrix [79]. Upon rehydration of water, the water content is approximately 80% and it is no more a matrix with jelly-like property. Similarly, the deprivation of water molecules by compression to form a thick, multilayered structure can be obtained [80]. Both cases result in a collagen matrix with much denser collagen concentration, which possesses visocelasticity.

The alternative method is the layer-by-layer deposition. This method is advantageous because the collagen aqueous solution is viscose and requires a long time to become a confluent solution when 2 collagen aqueous solutions of different concentrations are deposited on top of one another. The interface between the layers functions as a membrane, allowing water molecules from the sparse collagen layer to move to the dense layer, and the collagen concentration decreases near the interface. Then, the NaCl/Na2HPO4 salts diffuse into the collagen layers from the bottom, causing fibrillization toward the upper part of the matrix solution [81]. The resulting collagen matrix is one with a multilayered structure without clear boundaries. The cell infiltration differed according to the layers, where the cell infiltration was shown for the less dense side. This method is related to the complex formation as shown in Section 1.3.3 in Figure 1.3.

The integration of collagen matrices with different components was also developed. Tampieri et al. have reported that the collagen matrix with different hydroxyapatite concentration at each layer and hyaluronic acid on a specific layer can be prepared [82]. They have developed the layer compatible to cartilage on one side and bone on the other side. This implies that the collagen with different functionality at each side can be prepared. Similarly, Gillette et al. have reported the integration of the collagen-based fibril matrix. By increasing the temperature, the collagen solution and the collagen solution with alginate integrate to form a fibrillized collagen matrix with thick interface [83, 84]. This method does not require the layer-by-layer deposition and it shows the importance of collagen–collagen interface, where the actual bonding between the collagen matrices is controlled.

1.4 Conclusions and Future Perspectives

A lot of literature introduces various applications of collagen designed for biomaterials. However, there are not very many successful collagen-based products that can actually be found in the market. Moreover, scientific information on the physical, chemical and biological aspects of collagen and its behavior upon diverse treatment is still limited. The most difficult problem is lack of control of its physical, chemical and biological behavior. There are many reasons for these problems, but the most critical one is its special structure which is poorly understood. However, collagen still is an important material which challenges researchers. The knowledge about collagen is growing, and there were some very important breakthroughs over the last 10 years. We cannot ignore the good aspects of collagen such as its good biocompatibility and low antigenicity. Also, collagen is soluble in water and possesses functional groups which are relatively easy to chemically or physically modify in an aqueous condition. These benefits will encourage future developments and uses as indicated by the intensification of studies on the utilization of collagen in the growing fields of tissue engineering, regenerative medicine and drug delivery.

References

1. K. Fujioka, M Maeda, T. Hojo, A. Sano. Adv. Drug. Deliv. Rev. Vol. 31, p. 247, 1998.

2. T. Starborg, Y. Lu, R.S. Meadows, K.E. Kadler, D.F. Holmes. Methods. Vol. 45, p. 53, 2008.

3. H.K. Kelinman, D. Philp, M.P. Hoffman. Curr. Opin. Biotech. Vol. 14, p. 526, 2003.

4. W Friess. Eur. J. Pharma. Biopharma. Vol. 45, p. 113, 1998.

5. M.E. Nimni, D. Cheung, B. Strates, M. Kodama, K. Sheikh. J. Biomed. Mater. Res. Vol. 21, p. 741, 1987.

6. R. Zeeman, P.J. Dijkstra, P.B. van Wachem, M.J. van Luyn, M. Hendriks, P.T. Cahalan, J. Feijen. Biomaterials. Vol. 20, p. 921, 1999.

7. K. Yoshizato, A. Nishikawa, T. Taira. J. Cell. Sci. Vol. 91, p. 491, 1988.

8. K.S. Weadock, E.J. Miller, L.D. Bellincampi, J.P. Zawadsky, M.G. Dunn. J. Biomed. Mater. Res. Vol. 29, p. 1373, 1995.

9. N. Barbani, L. Lazzeri, C. Cristalli, M.G. Cascone, G. Polacco, G. Pizzirani. J. Appl. Polym. Sci. Vol. 72, p. 971, 1999.

10. L.H.H. Olde Damink, P.J. Dijkstra, M.J.A. van Luyn, P.B. van Wachem, P. Nieuwenhuis, J. Feijen. Biomaterials. Vol. 17, p. 765, 1996.

11. M.J.B. Wissink, R. Beernink, J.S. Pieper, A.A. Poot, G.H.M. Engbers, T. Beugeling, W.G. van Aken, J. Feijen. Biomaterials. Vol. 22, p.151, 2001.

12. M.J.A. van Luyn, P.B. van Wachem, L.H.H. Olde Damink. J. Biomed. Mater. Res. Vol. 26, p. 1091, 1992.

13. K. Nam, T. Kimura, A. Kishida. Macromol. Biosci. Vol. 8, p. 32, 2008.

14. K. Nam, T. Kimura, S. Funamoto, A. Kishida. Acta Biomater. Vol. 6, p. 403, 2010.

15. A.A.H.J. Sauren, M.C. van Hout, A.A. van Steenhoven, F.E. Veldpaus, J.D. Janssen. J. Biomech. Vol 16, p. 327, 1983.

16. F.J. Schoen, R.J. Levy. J. Biomed. Mater. Res. Vol. 47, p. 439, 1999.

17. R. Warren Sands, D.J. Mooney. Curr. Opin. Biotechnol. Vol. 18, p. 448, 2007.

18. E. Vorotnikova, D. McIntosh, A. Dewilde, J. Zhang, J.E. Reing, L. Zhang, K. Cordero, K. Bedelbaeva, D. Gourevitch, E. Heber-Katz, S.F. Badylak, S.J. Braunhut. Matrix Biol. Vol. 29, p. 690, 2010.

19. P. Bornstein, E.H. Sage. Curr. Opin. Cell Biol. Vol. 14, p. 608, 2002.

20. D. Barkan, J.E. Green, A.F. Chambers. Eur. J. Cancer. Vol. 46, p. 1181, 2010.

21. R.A. Allen, L.M. Seltz, H. Jiang, R.T. Kasick, T.L. Sellaro, S.F. Badylak, J.B. Ogilvie. Tissue Eng. Vol. 16A, p. 3363, 2010.

22. T.L. Sellaro, A, Ranade, D.M. Faulk, G.P. McCabe, K. Dorko, S.F. Badylak, S.C. Strom. Tissue Eng. Vol. 16A, p. 1075, 2010.

23. D.G. Attiah, R.A. Kopher, T.A. Desai. J. Mater, Sci. Mater. Med. Vol. 14, p 1005, 2003.

24. C.C. Xu, R.W. Chan, D.G. Weinberger, G. Efune, K.S. Pawlowski. J. Biomed. Mater. Res. Vol. 92A, p. 18, 2010.

25. J.E. Valentin, N.J. Turner, T.W. Gilbert, S.F. Badylak. Biomaterials. Vol. 31, p. 7475, 2010.

26. P.M. Crapo, T.W. Gilbert, S.F. Badylak. Biomaterials. Vol.32, p. 3233, 2011.

27. C. Yip. Ann. N.Y. Acad. Sci. Vol. 961, p. 109, 2002.

28. M.-T. Kasimir, E. Rieder, G. Seebacher, G. Silberhumer, E. Wolner, G. Weigel, P. Simon. Int. J. Artif. Organs. Vol. 265, p. 421, 2003.

29. S. Funamoto, K. Nam, T. Kimura, A. Murakoshi, Y. Hashimoto, K. Niwaya, S. Kitamura, T. Fujisato, A. Kishida. Biomaterials. Vol. 31, p. 3590, 2010.

30. Y. Hashimoto, S. Funamoto, S. Sasaki, T. Honda, S. Hattori, K. Nam, T. Kimura, M. Mochizuki, H. Kobayashi, A. Kishida. Biomaterials. Vol. 31, p. 3941, 2010.

31. C. Booth, S.A. Korossis, H.E. Wilcox, K.G. Watterson, J.N. Kearney, J. Fisher, E. Ingham. J. Heart Valve Dis. Vol. 11, p. 457, 2002.

32. F. Bolland, S. Korossis, S.P. Wilshaw, E. Ingham, J. Fisher, J.N. Kearney, J. Southgate. Biomaterials. Vol. 28, p. 1061, 2007.

33. K.E. Kadler, D.F. Holmes, J.A. Trotter, J.A. Chapman. Biochem. J. Vol. 316, p. 1, 1996.

34. J. Parkinson, K.E. Kadler, A. Bass. J. Mol. Biol. Vol. 247, p. 823, 1994.

35. D.J.S. Hulmes. Collagen Diversity, Synthesis and Assembly, Collagen: Structure and Mechanics, New York, 2008.

36. L. Fei, S. Perrett. Int. J. Mol. Sci. Vol. 10, p. 646, 2009.

37. N. Kuznetsova, S.L. Chi, S. Leikin. Biochemistry. Vol. 37, p. 11888, 1998.

38. E. Leikina, M.V. Mertts, N. Kuznetsova, S. Leikin. Proc. Natl. Acad. Sci. USA. Vol. 99, p. 1314, 2002.

39. V. Persikov, B. Brodsky. Proc. Natl. Acad. Sci. USA. Vol. 99, p. 1101, 2002.

40. J.M. Lee, D.R. Boughner, D.W. Courtman. J. Biomed. Mater. Res. Vol. 18, p. 79, 1984.

41. R. Zeeman, P.J. Dijkstra, P.B. van Wachem, M.J. van Luyn, M. Hendriks, P.T. Cahalan, J. Feijen. J. Biomed. Mater. Res. Vol. 51, p. 541, 2000.

42. K. Nam, T. Kimura, A. Kishida. Biomaterials. Vol. 28, p. 1, 2007.

43. K. Nam, Y. Sakai, S. Funamoto, T. Kimura, A. Kishida. J. Biomater. Sci. Polym. Edn. Vol. 22, p.1963, 2011.

44. B. Lanfer, U. Freudenberg, R. Zimmermann, D. Stamov, V. Körber, C. Werner. Biomaterials. Vol. 29, p. 3888, 2008.

45. M.-M. Giraud-Guille. J.Mol. Biol. Vol. 224, p. 861, 1992.

46. G. Mosser, A. Anglo, C. Helary, Y. Bouligand, M.-M. Giraud-Guille. Mat. Biol. Vol. 25, p. 3, 2006.

47. C. Helary, I. Bataille, A. Abed, C. Illoul, A. Anglo, L. Louedec, D. Letourneur, A. Meddahi-Pellé, M.-M. Giraud-Guille. Biomaterials. Vol. 31, p. 481, 2010.

48. G. Falini, S. Fermani, F. Foresti, B. Parma, K. Rubuni, M.C. Sidoti, N. Roveri. J. Mater. Chem. Vol. 14, p. 2297, 2004.

49. D.R. Henshaw, E. Attia, M. Bhargava, J.A. Hannafin. J. Ortho. Res. Vol. 24, p.481, 2006.

50. C. Guo, L.J. Kaufman. Biomaterials. Vol. 28, p. 1105, 2007.

51. J. Torbet, M.-C. Ronzière. Biochem. J. Vol. 219, p. 1057, 1984.

52. Y. Eguchi, M. Ogiue-Ikeda, S. Ueno. Neurosci. Lett. Vol. 351, p. 130, 2003.

53. A. Sionkowska, J. Skopinska-Wisniewska, M. Wisniewski. J. Mol. Liq. Vol. 145, p.135, 2009.

54. D.A. Wahl, J.T. Czernuszka. Eur. Cell Mater. Vol.11, p. 43, 2006.

55. M. Gelinsky, M. Eckert, F. Despang. Int. J. Mater. Res. Vol. 98, p. 749, 2007.

56. E. Ruoslahti, E. Engvall. Biochim. Biophys. Acta. Vol. 631, p. 350, 1980.

57. R.C. Cabrera, J.W. Siebert, Y. Eidelman, L.I. Gold, M.T. Longaker, H.G. Garg. Biochem. Mol. Biol. Int. Vol. 37, p. 151, 1995.

58. K. Nam, T. Kimura, A. Kishida, Biomaterials. Vol.28, p. 3153, 2007.

59. T. Nezu, F.M. Winnik. Biomaterials. Vol. 21, p. 415, 2000.

60. X. Duan, H. Sheardown. J. Biomed. Mater, Res. Vol. 75A, p. 510, 2005.

61. Z. Chen, X. Mo, F. Qing. Mater. Lett. Vol. 61, p. 3490, 2007.

62. T. Segura, P.H. Chung, L.D. Shea. Biomaterials. Vol. 26, p. 1575, 2005.

63. Y.-H Kim, C.-S. Cho, I.-K. Kang, S.Y. Kim, O.H. Kwon. Key Eng. Mater. Vol. 342–343, p. 101, 2007.

64. S.A. Sell, M.J. McClure, K. Garg, P.S. Wolfe, G.L. Bowlin. Adv. Drug Delivery Rev. Vol. 61, p. 1007, 2009.

65. D.I. Zeugolis, ST. Khew, E.S.Y. Yew, A.K. Ekaputra, Y.W. Tong, L.-Y.L. Yung, D.W. Hutmacher, C. Sheppard, M. Raghunath. Biomaterials. Vol. 29, p. 2293, 2008.

66. B. Brodsky, G. Thiagarajan, B. Madhan, K. Kar. Biopolymer. Vol. 89, p. 345, 2008.

67. H.M. Powell, S.T. Boyce. Tissue Eng. Vol. 15A, p. 2177, 2009.

68. M.G. Dunn, L.D. Bellincampi, A.J. Tria, Jr., J.P. Zawadsky. J. Appl. Polym. Sci. Vol. 63, p. 1423, 1997.

69. C.M. Haslauer, A.K. Moghe, J.A. Osborne, B.S. Gupta, E.G. Loboa. J. Biomater. Sci. Polym. Ed. Vol. 22, p. 1695, 2011.

70. T. Sato, G. Chen, T. Ushida, T. Ishii, N. Ochiai, T. Tateishi, J. Tanaka. Mater. Sci. Eng. Vol. 24 C, p. 365, 2004.

71. A. Sionkowska, M. Wisniewski, J. Skopinska. Polym. Degra. Stabil. Vol. 83, p. 117, 2004.

72. A. Sionkowska, A. Planecka, J. Kozlowska, J. Skopinska-Wisniewska. J. Mol. Liq. Vol. 144, p. 71, 2009.

73. A. Sionkowska. Progress Polym. Sci. Vol. 36, p. 1254, 2011.

74. E.D. Boland, J.A. Matthews, K.J. Pawlowski, D.G. Simpson, G.E. Wnek, G.L. Bowlin. Front. Biosci. Vol. 9, p. 1422, 2004.

75. L. Buttafoco, N.G. Kolkman, P. Engbers-Buijtenhuijs, A.A. Poot, P.J. Dijkstra, I. Vermes, J. Feijen. Biomaterials. Vol. 27, p. 724, 2006.

76. K. Nam, Y. Sakai, S. Funamoto, T. Kimura, A. Kishida. J. Biomater. Sci. Polym. Edn. Vol. 22, p.1963, 2011.

77. C.A. Miles, M. Ghelashvili. Biophys. J. Vol. 76, p. 3243, 1999.

78. C.A. Miles, N.C. Avery, V.V. Rodin, A.J. Bailey. J. Mol. Biol. Vol. 346, p. 551, 2005.

79. J. Engel, D.J. Prockop. Matrix Biol. Vol. 17, p. 679, 1998.

80. E.A. Abou Neel, U. Cheema, J.C. Knowles, R.A. Brown, S.N. Nazah. Soft Matter. Vol. 2, p. 986, 2006.

81. K. Nam, Y. Sakai, S. Funamoto, T. Kimura, A. Kishida. Soft Matter. In print.

82. A. Tampieri, M. Sandri, E. Landi, D. Pressato, S. Francioli, R. Quarto, I. Martin. Biomaterials. Vol. 29, p. 3539, 2008.

83. B.M. Gillette, J.A. Jensen, B. Tang, G.J. Yang, A. Bazargan-Lari, M. Zhong, S.K. Sia. Nature Materials. Vol. 7, p. 636, 2008.

84. B.M. Gillette, N.S. Rossen, N. Das, D. Leong, M. Wang, A. Dugar, S.K. Sia. Biomaterials. Vol. 32, p. 8067, 2011.

Chapter 2

Biological and Medical Significance of Nanodimensional and Nanocrystalline Calcium Orthophosphates

Sergey V. Dorozhkin

Abstract

Nano-sized particles and crystals play an important role in the formation of calcified tissues of various animals. For example, nano-sized and nanocrystalline calcium orthophosphates in the form of apatites of biological origin represent the basic inorganic building blocks of bones and teeth of mammals. Namely, according to recent developments in biomineralization, tens to hundreds of nanodimensional crystals of a biological apatite are self-assembled into these complex structures. This process occurs under a strict control by bioorganic matrices. Furthermore, both a greater viability and a better proliferation of various types of cells have been detected on smaller crystals of calcium orthophosphates. Thus, the nano-sized and nanocrystalline forms of calcium orthophosphates have a great potential to revolutionize the hard-tissue engineering field, starting from bone repair and augmentation to controlled drug delivery systems. This chapter reports on current state of the art and recent developments on the subject, starting from synthesis and characterization to biomedical and clinical applications. Furthermore, the chapter also discusses possible directions for future research and development.

Keywords: Calcium orthophosphates, hydroxyapatite, nanodimensional, nano-sized, nanocrystalline, biomedical applications, bone grafts, tissue engineering

2.1 Introduction

Living organisms can create amazing ways to produce various high-performance materials and over 60 different inorganic minerals of biological origin have already been revealed [1]. Among them, calcium orthophosphates are of special importance since they are the most important inorganic constituents of hard tissues in vertebrates [2, 3]. In the form of a poor crystalline, non-stoichiometric, ion-substituted CDHA (commonly referred to as biological apatite), calcium orthophosphates are present in bones, teeth, deer antlers and tendons of mammals to give these organs stability, hardness and function [2, 4, 5]. Though we still do not exactly know why highly intelligent animals use conformable calcium orthophosphates as their crucial biomineral for survival [6], current biomedical questions of persistent pathological and physiological mineralization in the body force people to focus on the processes, including the occurrence, formation and degradation of calcium orthophosphates in living organisms [7, 8, 9].

Biological mineralization (or biomineralization) is a process of in vivo formation of inorganic minerals [1, 2]. In the biomineralization processes, organized assemblies of organic macromolecules regulate nucleation, growth, morphology and assembly of inorganic crystals. Biologically formed calcium orthophosphates (biological apatite) are always nanodimensional and nanocrystalline, which have been formed in vivo under mild conditions. According to many reports, dimensions of biological apatite in the calcified tissues always possess a range of a few to hundreds of nanometers with the smallest building blocks on the nanometer size scale [2, 4, 5, 10, 11]. For example, tens to hundreds of nanometer-sized apatite crystals in a collagen matrix are combined into self-assembled structures during bone and teeth formation [2, 4, 5]. Recent advances suggest that this is a natural selection, since the nanostructured materials provide a better capability for the specific interactions with proteins [12].

Due to the aforementioned, nanodimensional and nanocrystalline forms of calcium orthophosphates are able to mimic both the composition and dimensions of constituent components of the calcified tissues. Thus, they can be utilized in biomineralization and as biomaterials due to the excellent biocompatibility [13, 14]. Further development of calcium orthophosphate-based biomaterials obviously will stand to benefit mostly from nanotechnology [15], which offers unique approaches to overcome shortcomings of many conventional materials. For example, nano-sized ceramics can exhibit significant ductility before failure contributed by the grain-boundary phase. Namely, already in 1987, Karch et al. reported that, with nanodimensional grains, a brittle ceramic could permit a large plastic strain up to 100% [16]. In addition, nanostructured ceramics can be sintered at lower temperatures; thereby major problems associated with a high temperature sintering are also decreased. Thus, nanodimensional and nanocrystalline forms of bioceramics clearly represent a promising class of orthopedic and dental implant formulations with improved biological and biomechanical properties [17].

Many other advances have been made in the biomaterial field due to a rapid growth of nanotechnology [18]. For example, a recent theory of aggregation-based crystal growth [19] and a new concept of mesocrystals [20, 21] highlighted the roles of nano-sized particles in biological crystal engineering. In this aspect, the study of calcium orthophosphates is a specific area in nanotechnology, because they might be applied readily to repair hard skeletal tissues of mammals [22–24].

Herein, an overview of nanodimensional and nanocrystalline apatites and other calcium orthophosphates in studies on biomineralization and biomaterials is given. The available calcium orthophosphates are listed in Table 2.1. To narrow the subject of this chapter, with a few important exceptions, undoped and unsubstituted calcium orthophosphates are considered and discussed only. The readers interested in various nanodimensional and nanocrystalline ion-substituted calcium orthophosphates [25–63] are referred to the original publications. Furthermore, details on calcium orthophosphate-based nanodimensional biocomposites [64–85] or nanodimensional calcium orthophosphate-based biocomposites [86–104] are available in references [105, 106].

Table 2.1 Existing calcium orthophosphates and their major properties [204, 205].

This chapter is organized into several sections. After a brief introduction (current section), general information on nano is provided in Section 2.2. Section 2.3 briefly compares the micron-sized and nanodimensional calcium orthophosphates. Section 2.4 briefly discusses the presence of nano-sized and nanocrystalline calcium orthophosphates in normal calcified tissues of mammals. The structure of nano-sized and nanocrystalline apatites is described in Section 2.5. Synthesis of nanodimensional and nanocrystalline calcium orthophosphates of various dimensions and shapes is reviewed in Section 2.6, while the biomedical applications are examined in Section 2.7. Finally, the summary and reasonable future perspectives in this active research area are given in Section 2.8.

2.2 General Information on Nano

The prefix nano specifically means a measure of 10−9 units. Although it is widely accepted that the prefix nano specifically refers to 10−9 units, in the context of nano-sized and nanocrystalline materials, the units should only be those of dimensions, rather than of any other unit of the scientific measurements. Besides, for practical purposes, it appears to be unrealistic to consider the prefix nano to solely and precisely refer to 10−9 m, just as it is not considered that micro specifically and solely concerns something with a dimension of precisely 10−6 m [107]. Currently, there is a general agreement that the subject of nanoscience and nanotechnology started after the famous talk: There’s plenty of room at the bottom, given by the Nobel Prize winner in physics Prof. Richard P. Feynman on December 26, 1959 at the annual meeting of the American Physical Society held at California Institute of Technology. This well-known talk has been widely published in various media (e.g., [108]).

In a recent extensive discussion about a framework for definitions presented to the European Commission, the nano-scale has been defined as being of the order of 100 nm or less. Similarly, a nanomaterial has been defined as any form of a material that is composed of discrete functional parts, many of which have one or more dimensions of the order of 100 nm or less [109]. Other definitions logically follow this approach such as: a nanocrystalline material is a material that is comprised of many crystals, the majority of which have one or more dimensions of the order of 100 nm or less (normally, with presence of neither the micron-sized crystals nor an intergranular amorphous phase), and a nanocomposite is a multi-phase material in which the majority of the dispersed phase components have one or more dimensions of the order of 100 nm or less [107]. Similarly, nanostructured materials are defined as the materials containing structural elements (e.g., clusters, crystallites or molecules) with dimensions in the 1–100 nm range [110], nanocoatings represent individual layers or multilayer surface coatings of 1–100 nm thick, nanopowders are extremely fine powders with an average particle size in the range of 1–100 nm and nanofibers are the fibers with a diameter within 1–100 nm [111, 112]. It also has been proposed to extend the lower size limit to 0.1 nm [113], which would include all existing organic molecules, allowing chemists to rightly claim they have been working on nanotechnology for very many years [114].

Strictly speaking, there are serious doubts that the term nanomaterial has a reasonable meaning. For example, let me cite Prof. David F. Williams, the Editor-in-Chief of Biomaterials: … some words which have no rational basis whatsoever become part of everyday language so rapidly, even if so illogically, that it is impossible to reverse the process and their common use has to be accepted, or perhaps, accommodated. Nanomaterial is one such word, where I have argued that it should not exist, but accept that it does through common usage and have to recognise its existence [107]. The discussion about nanomaterial provides a hint of the analysis of a biomaterial that follows, since a prefix, which is an indicator of scale, cannot specify the integer that follows (in this case a material) unless that integer can be qualified by that scale. In other words, it is very clear what a nanomaterial is because nano means 10−9 and a material is a measure of length. In the case of nanomaterial, what is it about the material that is 10−9. Is it the dimension of a crystal within the material, or of a grain boundary, a domain, or a molecule, or is it a parameter of a surface feature of the sample, or perhaps of the resistivity or thermal conductivity of the material. Clearly this is nonsense, but one has to accept that nanomaterials are here to stay, with even some journal titles containing the word. ([115], p. 5898, left column). Following this logic, such terms as nanocomposite, nanocoatings, nanopowders, nanofibers and nanocrystals are also senseless and should be replaced, for example, by composites with nano-sized (or nanodimensional) dispersed phase(s), coatings of nano-sized (or nanodimensional) thickness, nano-sized (or nanodimensional) powders, fibers of nano-sized (or nanodimensional) thickness and nano-sized (or nanodimensional) crystals, respectively. At least this has been done in this chapter.

According to their geometry, all nanodimensional materials can be divided into three major categories: equiaxed, one dimensional (or fibrous) and two dimensional (or lamellar) forms. Selected examples and typical applications of each category of nanodimensional materials and their use in biomedical applications are available in literature [116]. It is important to note, that in literature on calcium orthophosphates there are cases, when the prefix nano has been applied for the structures, with the minimum dimensions exceeding 100 nm [44, 83, 117–133].

As a rule, nanodimensional materials can be manufactured from nearly any substance. Of crucial importance, there are two major characteristics conferring the special properties of any nanodimensional material. These are the quantum effects associated with the very small dimensions (currently, this is not applicable to the biomaterials field) and a large surface-to-volume ratio that is encountered at these dimensions. For instance, specific surface areas for submicron-sized particles are typically 60–80 m²/g, while decreasing particle diameter to tens of nanometers increases the specific surface area up to 5 times more – an amazing amount of surface area per mass! Furthermore, all nanophase materials have unique surface properties, such as an increased number of grain boundaries and defects on the surface, huge surface area and altered electronic structure, if compared to the micron-sized materials [107, 134]. While less than ~ 1% of a micron-sized particle’s atoms occupy the surface positions, over a tenth of the atoms in a 10-nm diameter particle reside on its surface and ~ 60% in a 2-nm particle [135]. This very high surface-to-volume ratio of nanodimensional materials provides a tremendous driving force for diffusion, especially at elevated temperatures, as well as causes a self-aggregation into larger particles. Besides, solubility of many substances increases with particle size decreasing [136, 137]. What’s more, nanophase materials could have surface features (e.g., a higher amount of nano-scale pores) to influence the type and amount of adsorption of selective proteins that could enhance specific osteoblast adhesion [138]. Finally and yet importantly, the nanodimensional and nanocrystalline materials have different mechanical, electrical, magnetic and optical properties if compared to the larger grained materials of the same chemical composition [139–142].

The nanostructured materials can take the form of powders, dispersions, coatings or bulk materials. In general, nanostructured materials contain a large volume fraction (greater than 50%) of defects such as grain boundaries, interphase boundaries and dislocations, which strongly influences their chemical and physical properties. The great advantages of nanostructuring were first understood in the electronic industry with the advent of thin film deposition processes. Other application areas have followed. For example, nanostructured bioceramics were found to improve friction and wear problems associated with joint replacement components because they were tougher and stronger than coarser-grained bioceramics [143]. Furthermore, nanostructuring has allowed chemical homogeneity and structural uniformity to an extent, which was once thought to be impossible to achieve [110]. In calcium orthophosphate bioceramics, the major target of nanostructuring is to mimic the architecture of bones and teeth [144, 145].

2.3 Micron- and Submicron-Sized Calcium Orthophosphates versus the Nanodimensional Ones

The micron-sized calcium orthophosphate-based bioceramic powders suffer from poor sinterability, mainly due to a low surface area (typically 2–5 m²/g), while the specific surface area of nanodimensional calcium orthophosphates exceeds 100 m²/g [146]. In addition, the resorption process of synthetic micron-sized calcium orthophosphates was found to be quite different from that of bone mineral [147].

Although the nanodimensional and nanocrystalline features of natural calcium orthophosphates of bones and teeth had been known earlier [2, 148–153], the history of the systematic investigations of this field was started only in 1994. Namely, a careful search in scientific databases using various combinations of keywords nano + calcium phosphate, nano + apatite, nano + hydroxyapatite, etc., in the article title revealed 5 papers published in 1994 [154–158]. No papers published before 1994 with the aforementioned keywords in the title have been found.

Nanodimensional (size ~ 67 nm) HA was found to have a higher surface roughness of 17 nm if compared to 10 nm for the submicron-sized (~ 180 nm) HA, while the contact angles (a quantitative measure of the wetting of a solid by a liquid) were significantly lower for nano-sized HA (6.1) if compared to the submicron-sized HA (11.51). Additionally, the diameter of individual pores in nanodimensional HA compacts is several times smaller (pore diameter ~ 6.6 Å) than that in the submicron grain-sized HA compacts (pore diameter within 19.8–31.0 Å) [159]. A surface roughness is known to enhance the osteoblast functions while a porous structure improves the osteoinduction compared with smooth surfaces and nonporpous structure, respectively [138]. Furthermore, nanophase HA appeared to have ~ 11% more proteins of fetal bovine serum adsorbed per 1 cm² than submicron-sized HA [160]. Interfacial interactions between calcined HA nano-sized crystals and various substrates were studied and the bonding strength appeared to be influenced not only by the nature of functional groups on the substrate but also by matching of surface roughness between the nano-sized crystals and the substrate [161]. More to the point, incorporating nanodimensional particles of HA into polyacrylonitrile fibers was found to result in their crystallinity degree rising by about 5% [162]. In a comparative study on the influence of incorporated micron-sized and nano-sized HA particles into poly-L-lactide matrices, addition of nanosized HA was found to influence both thermal and dynamic mechanical properties in greater extents [163].

In general, nanostructured biomaterials [164] offer much improved performances over their larger particle-sized counterparts due to their huge surface-to-volume ratio and unusual chemical synergistic effects. Such nanostructured systems constitute a bridge between single molecules and bulk material systems [165]. For instance, powders of nanocrystalline apatites [166–172] and β-TCP [173] were found to exhibit an improved sinterability and enhanced densification due to a greater surface area. This is explained by the fact that the distances of material transport during the sintering becomes shorter for ultra-fine powders with a high specific surface area, resulting in a densification at a low temperature. Therefore, due to low grain growth rates, a low-temperature sintering appears to be effective to produce fine-grained apatite bioceramics [174]. Furthermore, the mechanical properties (namely, hardness and toughness) of HA bioceramics appeared to increase as the grain size decreased from sub-micrometers to nanometers [175].

More to the point, nano-sized HA is also expected to have a better bioactivity than coarser crystals [176–178]. Namely, Kim et al. found that osteoblasts (bone-forming cells) attached to the nano-sized HA/gelatin biocomposites to a significantly higher degree than to micrometer size analog [179]. An increased osteoblast and decreased fibroblast (fibrous tissue-forming cells) adhesion on nanophase ceramics [180–184], as well as on nanocrystalline HA coatings on titanium, if compared to traditionally used plasma-sprayed HA coatings, was also discovered by other researchers [185–187]. Scientists also observed enhanced osteoclast (bone-resorbing cells) functions to show healthy remodeling of bone at the simulated implant surface [177]. Besides, the proliferation and osteogenic differentiation of periodontal ligament cells were found to be promoted when a nanophase HA was used, if compared to dense HA bioceramics [188]. Thus, the underlying material property, responsible for this enhanced osteoblast function, is the surface roughness of the nanostructured surface [18]. Interestingly, an increased osteoblast adhesion was discovered on nano-sized calcium orthophosphate powders with higher Ca/P ratios [189], which points out some advantages of apatites over other calcium orthophosphates. Furthermore, a histological analysis revealed a superior biocompatibility and osteointegration of bone graft substitutes when nano-sized HA was employed in biocomposites [190–192]. However, data are available that nano-sized HA could inhibit growth of osteoblasts in a dose-dependent manner [193]. Furthermore, a cellular activity appeared to be affected by the shape and dimensions of nano-sized HA. Namely, the cellular activity of L929 mouse fibroblasts on nano-sized fibers with a diameter within 50–100 nm was significantly enhanced relative to that on a flat HA surface, while nanodimensional HA needles and sheets with a diameter/thickness of less than 30 nm inhibited cellular adhesion and/or subsequent activity because cells could not form focal adhesions of sufficient size [194].

Obviously, the volume fraction of grain boundaries in nanodimensional calcium orthophosphates is increased significantly leading to improved osteoblast adhesion, proliferation and mineralization. Therefore, a composition of these biomaterials at the nano-scale emulates the bone’s hierarchic organization, to initiate the growth of an apatite layer and to allow for the cellular and tissue response of bone remodeling. These examples emphasize that nanophase materials deserve more attention in improving orthopedic implant failure rates. However, to reduce surface energy, all nano-sized materials tend to agglomerate and, to avoid self-aggregation of calcium orthophosphate nanosized particles [195–198], special precautions might be necessary [54, 60, 120, 199–202].

Finally yet importantly, nano-sized crystals of CDHA obtained by precipitation methods in aqueous solutions were shown to exhibit physico-chemical characteristics rather similar to those of bone apatite [203]. In particular, their chemical composition departs from stoichiometry by calcium and hydroxide ions deficiency, leading to an increased solubility, and in turn bioresorption rate in vivo [148, 204–206]. The nano-sized crystals of CDHA also have a property to evolve in solution (maturation) like bone crystals. Namely, freshly precipitated CDHA has been shown to be analogous to embryonic bone mineral crystals whereas aged precipitates resemble bone crystals of old vertebrates [203].

2.4 Nanodimensional and Nanocrystalline Calcium Orthophosphates in Calcified Tissues of Mammals

2.4.1 Bones

Bone is the most typical calcified tissue of mammals and it comes in all sorts of shapes and sizes in order to achieve various functions of protection and mechanical support for the body. The major inorganic component of bone mineral is a biological apatite, which might be defined as a poorly crystalline, non-stoichiometric and ion substituted CDHA [2–5, 204–207]. From the material point of view, bone can be considered as an assembly of distinct levels of